The human eye functions to provide vision by transmitting light through a clear outer portion called the cornea, and focusing the image by way of a crystalline lens onto a retina. The quality of the focused image depends on many factors including the size and shape of the eye, and the transparency of the cornea and the lens.
When age or disease causes the lens to become less transparent, vision deteriorates because of the diminished light which can be transmitted to the retina. This deficiency in the lens of the eye is medically known as a cataract. An accepted treatment for this condition is surgical removal of the lens and replacement of the lens function by an artificial intraocular lens (IOL).
Intraocular lenses are employed as replacements for the crystalline lens after either extracapsular or intracapsular surgery for the removal of a cataract. In the United States, the majority of cataractous lenses are removed by a surgical technique called phacoemulsification. During this procedure, an opening is made in the anterior capsule and a thin phacoemulsification cutting tip is inserted into the diseased lens and vibrated ultrasonically. The vibrating cutting tip liquefies or emulsifies the lens so that the lens may be aspirated out of the eye. The diseased lens, once removed, is replaced by an artificial lens.
Intraocular lenses are generally of two types, those that are placed in the anterior chamber, i.e., between the iris and the cornea, and those that are placed in the posterior chamber, i.e., behind the iris. Both types of lenses are conventionally employed with the choice between an anterior chamber and a posterior chamber lens being partly dictated by requirements of the patient and partly dictated by the preferences of the physician inserting the lens. A third type of lens, known as iris-fixated lenses because they are secured to the iris periphery, can be thought of as being within one of the two types above, in that their optic portion is in either the anterior or posterior chamber.
Intraocular lenses normally consist of an optic with at least one and preferably two or more haptics that extend generally radially from the optic and contain distal portions that normally seat in the scleral spur for an anterior chamber lens and either in the ciliary sulcus or within the lens capsule for a posterior chamber lens. The optic normally comprises a circular transparent optical lens. The haptic in most lenses is a flexible fiber or filament having a proximate end affixed to the lens and having a distal end extending radially away from the periphery of the lens to form a seating foot. Several haptic designs are currently in use, for example, a pair of C-shaped loops in which both ends of each loop are connected to the lens, and, for example, J-shaped loops in which only one end of the loop is affixed to the lens.
Haptics are usually radially resilient and extend outwardly from the periphery of the lens and gently, but elastically, engage appropriate circumferential eye structures adjacent the iris or within the capsular bag. This resiliency is due to the conventional elastic properties of the materials of the haptic. The result is a haptic which when compressed and released will uncontrollably spring back immediately. This property makes the process of implantation and final positioning of the lens difficult since the haptics must be constrained during implantation. Also, once situated, the flexibility of the conventional haptic material makes the lens susceptible to decentration from being pushed by vitreous pressure from behind the lens or shifting due to pressure from adjacent ocular tissue. Also, the forces generated by the elastic recoil of the haptic release may damage the delicate local tissue.
The optimum position for a posterior chamber lens is in the capsular bag. This is an extremely difficult maneuver for the surgeon to accomplish. When a posterior chamber lens is employed it must be placed through the small pupillary opening, and the final haptic position is hidden behind the iris and not visible to the surgeon. It is therefore highly desirable to keep the overall dimensions of the posterior chamber lens as small as possible during implantation, letting it expand when it is finally situated where the surgeon intends, usually in the capsular bag. A small device is easier to manipulate in the eye, reduces the chance of the haptics coming in contact with the corneal endothelial tissue, and allows the surgeon ease of insertion, as he must often insert a lens with a 14 mm overall dimension through a pupil of 5 to 8 mm diameter. A smaller lens also reduces the lens/iris contact and can better guarantee that the intraocular lens and its haptics will be in the capsular bag.
In recent years intraocular lenses with and without haptics having relatively soft body portions have been provided such that the body portion could be folded generally across the diameter thereof for insertion into a smaller opening during implantation of the lens. Lenses formed of liquid or hydrogel constrained within a sheath have been designed which allow the lens body to be folded before insertion and then subsequently filled when in position. Unfortunately, the soft materials used for the bodies of these lenses lack the restorative strength sometimes required to return to their original shape.
Further, these lens types are typically deployed using either an elastic release mechanism, wherein mechanical energy stored by bending the elastic material is released when the mechanical constraint is removed, or through water uptake, also known as hydration, wherein the lens gradually absorbs water through an osmotic diffusion process. Both processes are difficult to control. In the former case, the elastic recoil may damage local tissue or may move the lens away from the center. In the latter case, the ultimate shape of the lens may become distorted if the expanding lens comes into contact with surrounding tissue. Further, hydrating materials are known to possess poor shape recovery properties.
In the natural lens, bifocality of distance and near vision is provided by a mechanism known as accommodation. The natural lens, early in life, is soft and contained within the capsular bag. The bag is suspended from the ciliary muscle by the zonules. Relaxation of the ciliary muscle tightens the zonules, and stretches the capsular bag. As a result, the natural lens tends to flatten. Tightening of the ciliary muscle relaxes the tension on the zonules, allowing the capsular bag and the natural lens to assume a more rounded shape. In this way, the natural lens can be focused alternatively on near and far objects. As the lens ages, it also becomes harder and is less able to change shape in reaction to the tightening of the ciliary muscle. This makes it harder for the lens to focus on near objects—a medical condition known as presbyopia. Presbyopia affects nearly all adults over the age of 45 or 50.
Typically, when a cataract or other disease requires the removal of the natural lens and replacement with an artificial IOL, the IOL is a monofocal lens, requiring that the patient use a pair of spectacles or contact lenses for near vision. Some bifocal IOLs have been created, but are not been widely accepted. Some IOL designs are single optic lenses having flexible haptics that allow the optic to move forward and backward in reaction to movement of the ciliary muscle. However, the amount of movement of the optic in these single-lens systems may be insufficient to allow for a useful range of accommodation. In addition, the eye must be medicated for one to two weeks to decrease eye movement in order for capsular fibrosis to entrap the lens that thereby provide for a rigid association between the lens and the capsular bag. Further, the commercial models of these lenses are made from a hydrogel or silicone material. Such materials are not resistive to the formation of posterior capsule opacification (“PCO”). The treatment for PCO is a capsulotomy using a Nd: YAG laser that vaporizes a portion of the posterior capsule. Such destruction of the posterior capsule may destroy the mechanism of accommodation of these lenses.
Known accommodative lenses also lack extended depth of focus in addition to having poor accommodation performance. Such known lenses further require precise lens sizing for proper function over a range of capsular bag sizes and lack long-term capsular fixation and stability. Further, as current lens replacement surgeries move towards smaller incision size, IOLs in general require the ability to be delivered through such small incisions.
Dual-optic lenses leverage the ability of the ciliary body-zonule complex to change the shape of the capsular bag. This allows the inter-lens distance to change, thereby allowing a change in refractive error. These dual-optic lenses can be large secondary to the optical hardware needed to create this optical system and requires larger corneal incisions to insert into the eye.
Intracorneal lenses are designed to treat refractive error or presbyopia. Intracorneal lenses include corneal implants and lenses, which are inserted through a small incision in the cornea created by a blade or a laser. The pocket formed by the incision in the cornea is used to position the implant to change the shape of the cornea. In the case of a lens implant, the pocket is used to position the refractive lens in the optically effective location. Some lenses create a pinhole-type effect to treat presbyopia. As current intraocorneal lenses move towards smaller incision size, devices in general require the ability to be delivered through such small incisions. Laser technology such as the femtosecond laser has enhanced the ability to create these smaller corneal wounds and pockets for implantation.
Phakic intraocular lenses are implanted either in the anterior chamber supported by the angle structures or in the posterior sulcus immediately posterior to the iris and anterior to the native lens. The lens is implanted through a minimally invasive wound at the limbus and inserted into or through the anterior chamber. The lenses are used to treat refractive error and have the risk of causing trauma to the lens and/or angle structures. Smaller incisions require folding the lens and then lens deployment in the eye, which increases the risk of damage to intraocular structures.
Known acrylic lens materials are unable to be compressed significantly to achieve desired functionality for IOLs. While various methodologies are known to fold or roll acrylic IOLs, these merely address the need to reduce the form factor of a deployed shape for the purposes of minimizing the required incision size for implantation. The actual volume displaced by these lenses remains constant so there is a limit on the minimum size that such IOLs can reach. Further, the ability to fold or roll these IOLs is limited by the ability of the material to resist strain caused by the stress of folding and return to a desired shape and provide the necessary optical qualities after implantation. Further, there is little control over the speed and force with which deployment of a lens occurs once it is implanted, which often causes trauma to tissues which engage haptics of the IOL.
The information included in this Background section of the specification, including any references cited herein and any description or discussion thereof is included for technical reference purposes only and is not to be regarded subject matter by which the scope of the invention as defined in the claims is to be bound.